1. Field of the Invention
The field of the invention is magnetic resonance imaging (MRI) and in particular local coils for use in receiving MRI signals.
2. Background Art
In MRI, a uniform magnetic field B.sub.0 is applied to an imaged object along the z axis of a Cartesian coordinate system, the origin of which is approximately centered within the imaged object. The effect of the magnetic field B.sub.0 is to align the object's nuclear spins along the z axis.
In response to a radio frequency (RF) excitation signal of the proper frequency, oriented within the x-y plane, the nuclei precess about the z-axis at their Larmor frequencies according to the following equation: EQU .omega.=.gamma.B.sub.0
where .omega. is the Larmor frequency, and .gamma. is the gyromagnetic ratio which is constant and a property of the particular nuclei.
Water, because of its relative abundance in biological tissue and the properties of its nuclei, is of principle concern in such imaging. The value of the gyromagnetic ratio .gamma. for water is 4.26 kHz/gauss and therefore in a 1.5 Tesla polarizing magnetic field B.sub.0 the resonant or Larmor frequency of water is approximately 63.9 MHz.
In a typical imaging sequence, the RF excitation signal is centered at the Larmor frequency .omega. and applied to the imaged object at the same time as a magnetic field gradient G.sub.z is applied. The gradient field G.sub.z causes only the nuclei in a slice through the object along a x-y plane to have the resonant frequency .omega. and to be excited into resonance.
After the excitation of the nuclei in this slice, magnetic field gradients are applied along the x and y axes. The gradient along the x axis, G.sub.x, causes the nuclei to precess at different frequencies depending on their position along the x axis, that is, G.sub.x spatially encodes the precessing nuclei by frequency. The y axis gradient, G.sub.y, is incremented through a series of values and encodes y position into the rate of change of phase of the precessing nuclei as a function of gradient amplitude, a process typically referred to as phase encoding.
A weak nuclear magnetic resonance generated by the precessing nuclei may be sensed by the RF coil and recorded as an NMR signal. From this NMR signal, a slice image may be derived according to well known reconstruction techniques. An overview NMR image reconstruction is contained in the book "Magnetic Resonance Imaging, Principles and Applications" by D. N. Kean and M. A. Smith.
The quality of the image produced by MRI techniques is dependent, in part, on the strength of the NMR signal received from the precessing nuclei. For this reason, it is known to use an independent RF receiving coil placed in close proximity to the region of interest of the imaged object to improve the strength of this received signal. Such coils are termed "local coils" or "surface coils". The smaller area of the local coil permits it to accurately focus on NMR signal from the region of interest. Further, such local coils may be of higher quality factor or "Q" than the RF transmitting coil increasing the selectivity of the local coil and the relative strength of the acquired NMR signal.
The smaller size of the local coil makes it important that the local coil be accurately positioned near the region of interest. For "whole volume" coils, employing two antenna loops to receive the NMR signal from a volume defined between the loops, accurate positioning of the coils is particularly important. For a whole volume neck coil, for example, the two antenna loops must be placed on opposite sides of the neck and yet generally opposed along a single axis. This may be accomplished by fitting the coils to the surface of a cylindrical form, the form parting along a plane intersecting the axis of the cylinder to form two halves. These halves may be clamped about the patient during the scan.
One problem with this method of mounting the coils is that the separation of the coils is fixed by the radius of the cylinder. Equally important, the joining of the two halves of the form creates a closed cage that may create a disquieting sense of confinement.
A major technical problem in NMR systems is "decoupling" the local coil from the RF excitation signal from the transmit coil during the application of the RF excitation signal. Such decoupling reduces the distortion of the excitation field by the local coil and prevents potential damage to the sensitive circuits connected to the local coil from possibly large induced voltages.
Inductive coupling between the excitation field and the local coil may focus the deposition of the RF energy on a reduced volume the imaged object. In the case of the medical imaging of a patient, such focused energy may cause burns. Energy coupled to the local coil itself may cause heating of that coil, producing indirect burns to the patient and damage to the local coil and its circuitry. The problem of distortion and inductive coupling is compounded by the typical high "Q" of the local coils.
One method of decoupling the local coil from the RF excitation field is through the use of one or more solid state switches positioned along the local coil which may be activated either by an external electrical signal (active decoupling) or by the RF excitation field itself (passive decoupling). These switches disable or detune the local coil. One such approach which shows generally the use of back-to-back diodes for passively decoupling a local coil is described in U.S. Pat. No. 4,725,779, issued Feb. 16, 1988 to Hyde et al., entitled: "NMR Local Coil with Improved Decoupling" and hereby incorporated by reference. In this reference, back-to-back diodes, in the presence of the large induced voltages from the transmit coil, short together two adjacent antenna coils having counter rotating currents thus decoupling the antenna coils from the RF excitation field and reducing the inductive coupling to the local coil. The advantage of passive decoupling is the elimination of the need for additional wires and signals to control the decoupling device and hence the simplification of the coil.